As more members of the “baby boomer” generation age, the number of surgical procedures to maintain a youthful appearance continues to increase. Of these cosmetic procedures, a significant increase has been in the area of non-invasive aesthetic applications. Many laser based systems are on the market with FDA clearance to non-ablatively treat wrinkles and rejuvenate skin texture. Lasers treat the skin and underlying subcutaneous tissues by depositing light energy to heat the tissues. The depth of treatment, however, is limited by the laser wavelength.
An alternative heating method is radiofrequency (RF) heating which provides variable heat penetration. RF energy can be delivered to skin tissues for aesthetic and therapeutic effect using either monopolar or bipolar electrode-coupled induction techniques. These systems require the use of active cooling at the interface between the skin surface and the electrodes to prevent localized burning.
The clinical opportunity in the field of skin tightening/wrinkle removal is therefore very significant and is an accepted treatment. However, that established market is being addressed to varying degrees by technologies which, though they are state-of-the-art today, leave the physician and customer base less than satisfied. These procedures for example require aggressive skin cooling in conjunction with laser/light therapies to provide treatment from just below the epidermis to approximately 0.5-1.0 mm below the skin surface, as well as the various RF induction methods, with low reproducibility of clinical results/outcomes due to inherent limitations in the physics of the approach.
It has long been known that damaging collagen will cause shrinkage and neocollagenesis (rejuvenation). It has been shown that the physiology will allow excellent clinical results that will allow physicians and device companies to serve patients profitably—but they have also proven that those results are often inconsistent. In some cases the clinical outcome is dramatic, in others it is imperceptible, and in still others the end result can be worse than the initial condition (significant burns, overshrinkage and loss of skin form). The cause of these inconsistent clinical results is related to the types of technology applications that apply the thermal energy with these devices—which are, themselves, inconsistent. The goal of the treatment is to heat the underlying tissue (dermis), and some of the deeper tissue, at temperatures ranging from 55° C. to 70° C. for a short period of time while leaving the surface (epidermis) and underlying tissue unaffected. Existing technologies are not able to accurately control where they apply the therapeutic treatment, depth of penetration, or how much therapeutic energy is absorbed by the target region.
There are three primary approaches being pursued by conventional systems using energy to treat skin and subcutaneous fat for aesthetics purposes, (1) disruption of the fat cells through agitation and cavitation to affect liposuction (may be performed invasively or noninvasively), (2) affecting thermal injury to the skin surface (epidermis) to stimulate neocollagenesis (the forming of new tissue) to smooth the texture of the skin, and (3) affecting thermal elevation of the tissue underlying the epidermis to affect removal of deeper wrinkles and tighten sagging skin.
Variations in High Intensity Ultrasound technologies—cellular disruption vs acoustic stimulation to heat tissue. To those who are not experts in the field it may appear that High Intensity Focused Ultrasound (HIFU) is the primary term used to describe the application of acoustic energy for thermal therapy applications and that there are several participants in this field. Actually, HIFU is specific to a particular method of delivery of acoustic energy, and does not encompass several other methods to use ultrasound for treatment.
There are five conventional variations to therapeutic applications of ultrasound:                (1) Low intensity, low frequency stimulation of bone tissue to encourage bone healing or to increase membrane permeability for the purpose of increased membrane transport of chemical agents.        (2) High intensity, low frequency application to affect cellular disruption. The primary applications for this family of devices are for disruption of fat cells in liposuction or disruption of thromoboses in vascular structures.        (3) Low intensity, high frequency application to affect therapeutic heating for muscle soreness. A variety of products in the field of sports medicine have been employed for years.        (4) High intensity, high frequency application to produce molecular agitation and directly interact with the high frequency mechanical properties of the tissue to produce localized heating within a desired therapeutic zone:                    a. The delivery approaches vary, and the use of hemispherical focused transducers is incorporated in the prior art products, and this is the typical HIFU. These include products for “spot” ablation of cancerous tissue and Benign Prostate Hyperplasia (BPH), creating cardiac lesions to treat atrial fibrulation (E), and tissue dissection/tissue welding.            b. Technology that uses tubular and curvilinear soft-focus and line focus transducer technology in both singular and array structures to create a customized shaped volume region of therapy. This can be achieved through explicit transducer design on an a priori basis and using multiple element designs integrated to permit dynamic adjustment of the therapeutic size and shape, dependent upon the specific tissue treated. Thus, volumetric heating of customized shapes and sizes can be achieved. For mid-size and larger regions, this permits treatment times that are much shorter than achieveable with “step focused” systems. Further, the control of the customized shape and treatment volume is exquisite, permitting an exact lesion size or treatment region to be created.                        (5) Acoustic Shock Wave Lithotripsy (ASWL) for disruption of calcium deposits such as kidney stones and bone spurs.        
Regarding methodology 4(a) above, (HIFU) approaches use hemispherical transducers to create focal points of energy (see FIG. 1). This approach works well when the desired result is to create a “cigar-shaped” lesion as the approach would produce a very high intensity energy density in the lateral cross section at the focal depth with a focal length of approximately eight times the lateral focal cross section which is centered at the focal depth. An example would be an external or intracavitary transducer focused at a depth of 3 cm that has a focal zone with a 1 mm cross section and a focal length of 8 to 10 mm. Depending upon the frequency, focal length, focal gain and input power, it is possible to create extremely high power densities at the center of each focal zone. Exquisite control of such energy using real-time, spatially-registered imaging is a requirement to deliver treatment that doesn't leave “gaps” laterally and doesn't seriously injure nearby normal tissues.
Creating a volumetric lesion with standard HIFU approaches would require the creation of multiple small lesions to cover the desired lateral cross section. As an example, a 1 cm2 square lateral region would require approximately eight half-power-width overlapping zones in both lateral directions, producing a 1 cm×1 cm lateral by 1 cm depth zone of temperature elevation. This would require the creation of 64 separate focal zones. Treatment using such an approach would be slow (approximately 60 seconds for a 1 cm region) and non-uniform in treatment.
When affecting a thermal increase in deeper tissue while leaving the tissue adjacent to the applicator probe (i.e. the skin) relatively unaffected, focused ultrasound technology is intrinsically superior to radiofrequency methods for two reasons:                (1) The electrical properties of various tissue types (epidermis, dermis, subcutaneous fat, fibrous septae, and subcutaneous muscle) vary much more than the acoustical properties of those tissue structures. This is because the electrical properties are dominated by water and electrolyte (salt) content, whereas the acoustic properties are predominately dependent on density differences. The result in this wider variation is that the tissue resistivity. Therefore, RF energy is not uniformly absorbed by the tissue below the application probe.                    RF power is not propagated through the tissue. RF is resistive in absorption, i.e. like connecting a network of resistors in a series-parallel combination across a big battery and heating the resistors along the available current pathways. Any propagation of the resultant heat is due to the thermal conductivity of the tissue. Any propagation of the resultant heat to neaby tissue is due to the thermal conductivity of the respective tissue. Small variations in tissue composition and variations in blood perfusion, therefore, can dramatically affect the electrical properties of the tissue and the energy absorption profile with RF treatment (and thus the treatment efficacy) of the underlying tissue. This phenomenon will be discussed in greater detail below.            With a more consistent energy absorption profile from energy that is propagated through the tissue (with ultrasound) the energy absorption (and treatment efficacy) are more uniform and predictable.                        (2) Because RF is a resistive heating phenomenon, dependent on the current density in the tissue, most of the RF effect occurs directly at the electrode/skin interface. Between 50% and 90% of the current (thus resistive heating) occurs in the 750 um of epidermis (a region which must be cooled to prevent skin burns). This means that most of the energy is dissipated and unproductive. Not only is this inefficient, but if there is a variation in tissue characteristics in the region within and below the cooled zone, dramatic changes in energy disposition to the region outside of the cooled zone could occur. Paths of high tissue conductivity next to those that are more resistive produce widely varying RF absorption patterns, often dramatically affecting resultant heating patterns.                    To illustrate this point further—if 75% of the energy is supposed to be dissipated in the cooling process, then only 25% of the energy is delivered to the region to be treated. If the low resistance components (saline from sweat, etc.) are twice as prevalent in the 750 um surface zone, then more energy (than expected) will be delivered to the deeper zone. Since there is no consistent means of monitoring where this energy is deployed, there could be rapid heating and tissue overtreatment in some areas and undertreatment in others within this region.                        
To illustrate the first point in more detail, the graph in FIG. 2 depicts the known electrical properties of several types of tissue. For example, fatty tissue has impedances ranging from 1,600 ohms to 2,000 ohms, while blood has a resistance of 150 ohms to 200 ohms. Since the cross sectional area of a small vessel could be 0.5 mm, such a structure in a 1.5 cm2 treatment zone represents only 0.5% of the area through which current would travel, but the current density would be ten times that of the surrounding tissue. Therefore, one structure, which is only 0.5% of the current-carrying media, would carry 3.5% of the current to the underlying tissue.
In conventional RF systems, such as described in: “Selective Fibrous Septae Heating”. An additional mechanism for Capacitively Coupled Monopolar Radiofrequency”, (Karl Pope, Mitch Levenson, E. Vic Ross, MD), the subcutaneous tissue is described as being a network of blood vessels and collagenous structures which connect the dermis to the underlying muscle (the fibrous septae). This anatomy is depicted in FIG. 3. It is asserted that for such RF systems the fibrous septae have significantly lower electrical impedance properties than the surrounding fatty tissue. An infrared image obtained during treatment (FIG. 4 from the prior art cited above) shows non-uniform heating of the underlying tissue. The probe 150 is visible and a cooled region 160 is directly below. The heated regions 170 appear non-uniform within the normal tissue 180. It should be emphasized that these are untouched replications from the literature article, and the lines (drawn by the authors of the paper) are estimations of the location of fibrous septae. The authors' intention was to depict deeper heating and contraction of the fibrous septae.
This image from the prior art literature provides significant clues of the shortcomings of the RF procedure. In FIG. 4 most of the heating appears to be at the fat/muscle interface—about 4 mm to 6 mm below the surface of the probe 150. This would make sense if, as the authors claim, the surface is being cooled and the excess current is being shunted through the surface tissue to the underlying muscle. The intense heating at this interface can explain the fat necrosis which has occurred in several RF heating cases known in the art. In fact, there is a zone in the fat/muscle interface depicted in the tissue cross-section, which appears to show some fat necrosis.
Note also that the heating under the probe 150 is not uniform. There are portions of intense heating in the epidermis (see the bright yellow region on the right, behind the label “Dermis”), while the cooling effect seems to occur under only half of the probe 150. The 2 mm of dermal tissue contains regions of the heated region 170 and the unheated normal tissue 180.
Another fact that the reader should note is that, even in the image depicting a prior art “shallow probe” (See FIG. 5), that there is very little cooling in the epidermal region immediately adjacent to the probe 150 and there is a zone of significant heating at the fat/muscle interface 190 some 4 mm to 6 mm deep.
The tissue cross-sectional photograph from the prior art presents some inconsistencies in the theory. FIG. 3 shows a dense network of fibrous septae—but the infrared (IR) photograph in FIG. 4 shows only a few shunting paths. Either the shunting is occurring along different paths (perhaps blood vessels), or some fibrous septae are more electrically conductive than others. Further, the tissue cross-section photograph shows no shrinkage in the fibrous septae-rich subcutaneous zone. There is significant shrinkage in the dermal layer—as one would expect—but there is none in the underlying zone. The thermal shunting in the underlying zone seems only to contribute to fat necrosis.
These images, indicate that uniform heating with RF is not easily achieved (if it is achievable at all). The acoustic properties of tissue are much more uniform, with the acoustic absorption between brain, kidney, liver, and muscle at a given frequency varying by 15% or less (for purposes of showing the state of the art, see Table 4.19, page 116 in Duck, F A, Physical Properties of Tissue, Academic Press, 1990. Tissue heating is a function of the acoustic velocity and attenuation through the specific tissue type. In general, absorption is directly related to the tissue density. Unlike the case for RF, acoustic energy actually is transmitted through soft tissues and it loses energy to heat conversion as it propagates. By selecting the frequency and focusing parameters judiciously, a large portion of the propagated energy is converted to heat directly in the desired region. The high degree of directivity is attained because at higher frequencies (in the MHz range) the wavelengths are short and can be directed and/or focused, just like light. However, the penetration is significantly greater. Insert Fat Necrosis Comment w/acoustic absorption.
Another conventional device employs a bipolar electrosurgical approach. The theory behind this approach is that the current would pass from one electrode to the other, staying in the underlying tissue. Unlike the RF electrodes described for the RF system, however, which are planar, the electrodes may be considered to approach two short linear sources. As such, the current density (and associated power) fall off as a factor of 1/R3 (as opposed to 1/R2 or 1/R for the planar approach. Most of the current flows along the surface of the tissue. It is virtually impossible to create any heating at depth.
Depending on whether the dissipation mechanism falls off at 1/R2 or 1/R, the monopolar approach results in between 50% and 90% of the energy is applied to the 750 um of epidermal tissue which is cooled to prevent burning (see FIG. 6A). The bipolar system of FIG. 6B, with a dissipation mechanism that falls off at 1/R3, more than 99% of the energy is applied to that 750 um cooled zone. This design feature leaves very little energy available for tissue treatment. Only 10% to 50% of the thermal effect occurs in the zone outside of the cooled epidermis (blue zone in graph) with the monopolar approach. With the bipolar approach of FIG. 6B only 1% of the thermal effect occurs in the zone outside the cooled epidermis. A significant quantity of thermal effect occurs beyond the therapeutic zone with the monopolar approach.
The above described electrosurgical methods for deep skin heating are not uniform and/or not predictable, or produce so little thermal action that they are ineffective. The bipolar linear electrodes produce very little effect. The monopolar planar electrodes allow current to shunt through low impedance structures to produce non-uniform heating in the dermis with a concentration at the fat/muscle interface, which could contribute to fat necrosis.
In light based treatment approaches, the theory behind light-based deep tissue heating requires applying a radiant energy source which dissipates as a function of depth while cooling the surface. The method of action is actually very different from the RF approach, but both have the result that the thermal effect is significantly greater at the probe/tissue interface (skin) than in deeper layers.
With light-based approaches, the molecular entities in the tissue (primarily water) absorb the photons from light-based energies and convert that energy (more-or-less) directly to heat, and that the light energy dissipates much less dramatically than RF. For instance, according to Franceschini, et al, (“Near-Infrared Absorption and Scattering Spectra of Tissues in Vivo”) presented at the SPIE in 1999 (http://www.eotc.tufts.edu/Documents/Faculty/Franceschini/papers/spie99-mari.PDF), the absorption rate of infrared light in skin tissue is approximately 20%/cm.
A chart of this absorption profile is presented in FIG. 7. Note that, although a significant amount of applied energy is transmitted beyond the cooled zone (85%), only 20% is absorbed in the therapeutic zone of the dermis, another 20% is absorbed in the subcutaneous fat, and more than 40% is transmitted into the deeper muscle tissue. 15% of the light is absorbed in the cooled zone of the epidermis, 20% in the therapeutic zone of the dermis, 20% in the subcutaneous fat, and 45% in the deeper muscle tissue. Although the skin is spared damage by the cooling process, there is no means of controlling thermal injury to the subcutaneous fat and deeper muscle.
The light-absorbing characteristics of tissue are much more uniform than the electrical characteristics. Thus, light absorption is more gradual and doesn't exhibit the large unpredictabilities found with RF approaches. In Franceschini's paper referenced above, the absorptive characteristics of the three patients ranged from 10%/cm to 25%/cm (a factor of 2.5, while the difference between the electrical impedance of fat and blood could be a factor of 10). There are factors, however, such as the concentration of melanocytes (such as with certain ethnic groups, or variations after recent exposure to the sun) that also affect the absorption levels. Melanocytes act as “absorbers” that selectively absorb light energy, producing inhomogeneous energy absorption, depending upon the amount degree of their presence and uniformity. This can yield highly variable results in such instances. Although the skin is spared thermal injury by the cooling process, there is no means of controlling injury to the subcutaneous fat and deeper muscle.
In another conventional device shown in FIG. 8, infrared energy is applied to the underlying tissue. The prior art device transmits infrared light in the range of 1,100 nm to 1,800 nm. The contact head has a cooling mechanism to protect the skin from burning. There is no mechanism to protect the deeper subcutaneous fat and muscle from thermal injury.
In another light based system (not shown), a NdYAG laser is used and, which transmits in the 1,064 nm range. The device also incorporates epidermal cooling to spare skin damage. This type of devices operates on the principle of applying infrared energy to the underlying tissue. The device transmits infrared light in the range of 1,100 nm to 1,800 nm. The contact head has a cooling mechanism to protect the skin from burning.
Light-based energy sources can effectively heat the near dermal and subcutaneous layers to affect treatment. They transmit only a small portion of their energy in these regions, however, and they cannot control the energy applied to deeper subcutaneous fat and muscle. In order to avoid injuring these deeper structures, they must limit the amount of energy applied altogether. This, in turn, limits the amount of energy applied to the dermal zone and thus, the effectiveness of the treatment. Further, the presence of variable degrees of melanocytes can produce unpredictable variability in absorption.